There have been broadly employed radiographic images such as X-ray images for diagnosis of the conditions of patients on the wards. Specifically, radiographic images using a intensifying-screen/film system have achieved enhancement of speed and image quality over its long history and are still used on the scene of medical treatment as an imaging system having high reliability and superior cost performance in combination. However, these image data are so-called analog image data, in which free image processing or instant image transfer cannot be realized.
Subsequently, there appeared computed radiography (also denoted simply as CR) as a radiographic image detection apparatus in a digital system. In the CR, digital radiographic images are directly obtained and can be displayed on an image display apparatus such as a cathode tube or liquid crystal panels, which renders it unnecessary to form images on photographic film and results in drastic improvement of convenience for diagnosis in hospitals or medical clinics.
The CR has been accepted mainly in medical sites, where X-ray images are obtained using a photostimulable phosphor plate. The photostimulable phosphor plate is one in which a radiation having been transmitted through an object is accumulated and excited in a time-series manner upon exposure to electromagnetic waves (exciting light) such as infrared light, whereby the accumulated radiation is emitted as stimulated emission at an intensity corresponding to the radiation dosage and which is constituted of a laminar photostimulable phosphor on a prescribed substrate.
However, this photostimulable phosphor plate, which is not sufficient in signal-to-noise ratio or sharpness and is also insufficient in spatial resolution, has not yet reached the image quality level of the conventional screen/film system.
Further, there appeared, as a digital X-ray imaging technology, an X-ray flat panel detector (FPD) using a thin film transistor (TFT), as set forth in, for example, an article “Amorphous Semiconductor Usher in Digital X-ray Imaging” described in Physics Today, November, 1997, page 24 and also in an article “Development of a High Resolution, Active Matrix, Flat-Panel Imager with Enhanced Fill Factor” described in SPIE, vol. 32, page 2 (1997).
The FPD has advantages such that it is superior to the CR in terms of downsizing of the apparatus being feasible and a moving image display. However, similarly to the CR, the FPT has not yet reached the image quality level of the screen/film system, so that desire for high image quality increased recently.
The FPD system employs a scintillator plate made of an emissive X-ray phosphor to convert radiation to visible light, in which electrical noise generated in TFT or in the circuit to drive the TFT is relatively high, so that even in imaging at a low dose, the SN ratio is lowered, making it difficult to ensure emission efficiency to maintain desired image quality level.
Generally, the emission efficiency of a scintillator plate depends of the phosphor layer thickness and X-ray absorbance of the phosphor. A thicker phosphor layer causes more scattering of emission within the phosphor layer, leading to deteriorated sharpness. Accordingly, necessary sharpness for desired image quality level necessarily determines the layer thickness.
Specifically, cesium iodide (CsI) used in the phosphor layer of a scintillator plate exhibits a relatively high conversion rate of from X-ray to visible light. Further, a columnar crystal structure of the phosphor can readily be formed through vapor deposition and its light guide effect inhibits scattering of emitted light within the crystal, enabling an increase of the phosphor layer thickness.
To achieve enhanced sharpness of a scintillator plate, there was proposed a manufacturing method of a radiation image conversion panel, comprising forming a phosphor layer through gas phase deposition. Gas phase deposition methods include a vapor deposition method and a spattering method. In the vapor deposition method, for instance, a vaporization source composed of raw phosphor material is heated by a resistance heater or exposure to an electron beam to vaporize the vaporization source and allow a distillate to be deposited on the substrate surface, forming a phosphor layer comprised of columnar phosphor crystals. The use of a single vaporization source makes it feasible to achieve uniform deposition.
In the formation of a phosphor layer, since the use of CsI alone results in lowered emission efficiency, there may be used various additives. It is known that an additive content of not less than 0.01 mol %, based on CsI enhances emission efficiency.
There was disclosed a technique for use as an X-ray phosphor in which a mixture of CsI and sodium iodide (NaI) at any mixing ratio was deposited on the substrate to form sodium-activated cesium iodide (CsI:Na), which was further subjected to annealing as a post-treatment to achieve enhanced visible-conversion efficiency, as set forth in JP-B No. 54-35060 (hereinafter, JP-B refers to Japanese Patent Publication).
Recently, there was also disclosed a technique for preparing an X-ray phosphor in which CsI is formed through deposition and activation material such as indium (In), thallium (Tl), lithium (Li), potassium (K), rubidium (Rb) or sodium (Na) was formed by spattering, as set forth in, for example, JP-A No. 2001-59899 (hereinafter, the term JP-A refers to Japanese Patent Application Publication).